It has been recognized that when heat is applied to the areas of animal or human tissue containing malignant cells so as to increase the temperature of the area to 41.degree.-45.degree. C. a preferable decrease in the viability of malignant cells occurs. The decreased viability of the malignant cells results either in the cell death or cell sensitization to the effects of concomitant chemotherapy and/or radiation therapy and therefore is favorable for the tumor therapy. Such heat treatment of tissues is known as hyperthermia [1,2]. In the group of hyperthermia techniques called ferromagnetic hyperthermia, a particulate ferromagnetic material is confined within the treatment area and further exposed to the oscillating electromagnetic field. The confined ferromagnetic particles dissipate the energy of the field in the form of heat through various kinds of energy losses and therefore cause hyperthermia in the area of their confinement [3-16].
A variety of ferromagnetic microparticulate and colloidal materials have been reported for the use in ferromagnetic hyperthermia including: oxidized pentacarbonyl iron particles [3-5], iron powders [6,7], magnetite dispersions [8,9], ceramic-magnetic iron composites [10,11], alkali earth hexaferrites [12], dextran-ferrite [13,14], and oxidized dextran-magnetite compounds [15,16]. The magnetic properties of such particles have not been disclosed in terms of standard tests which would permit direct comparison of different preparations. The heating quality of the above materials are compared herein on the basis of available literature data or by our own measurements of the materials prepared according to the reported procedures. The heating quality of the ferromagnetics are expressed herein as their specific power absorption rate (SAR) defined as the amount of heat released by a unit weight of the material per unit time during exposure to an oscillating magnetic field of defined frequency and field strength. Physical considerations of the mechanisms involved in the dissipation of the electromagnetic field energy by these materials suggest that SAR is approximately proportional to the field frequency times field strength squared. Therefore for fair comparison we have normalized the heating rates to the frequency of 1 MHz and peak field strength of 100 Oersted.
The results of such comparison (Table 1) suggested that all previously disclosed materials require substantial tissue concentrations (more than 5-10 mg/gram of tissue) to provide effective heating. Materials based on precipitated iron oxides generally gave superior results compared with other types of ferromagnetics. Precipitated iron oxides are well-known in such applications as magnetic resonance imaging in diagnostic medicine [17,20,21], magnetic labeling and sorting of cells [18], and bone marrow purging for treatment of cancer patients [19]. These compounds were shown to produce little, if any, toxicity to animals and humans [20,21]. It has also been shown that dextran-coated iron oxide particles have excellent stability against aggregation and precipitation in physiological media [16,22,23]. A widely used procedure for precipitating iron oxide in the presence of dextran is described by Molday [23]. According to the Molday procedure, dextran (stabilizer) is first mixed with ferric and ferrous salts in aqueous solution; then the excess of aqueous ammonia (precipitant) is added to precipitate iron hydroxides, and the suspension is heat treated to convert hydroxides into magnetic iron oxide. Finally, the grossly aggregated fraction is separated by low speed centrifugation, and the remaining colloid is purified by dialysis and/or gel-filtration. As described below, magnetic particles prepared according to the present invention provide substantially increased tissue heating compared to previously disclosed particles. Therefore the particles of the present invention substantially reduce the amount of magnetic material required to achieve therapeutically useful heating.
The Molday procedure [23] includes:
1) combining of iron(II) and iron(III) chloride salts with the stabilizer (dextran) and precipitant (ammonia) in aqueous solution with agitation;
2) treatment of the reaction mixture at the temperature of 60.degree.-65.degree. C. for 15 min;
3) separation of larger aggregates by low-speed centrifugation; and PA1 1. The particles are magnetic iron oxide particles associated with a polymer (polysaccharide). PA1 2. The particles are superparamagnetic, i.e., possess negligible coercivity. PA1 3. The particles form a suspension in a physiologically acceptable, water-based medium, and this suspension is stable against precipitation under gravity. PA1 4. The particles have average magnetic moment of at least about 10.sup.-15 erg/Gauss or more, up to about 1.6.times.10.sup.-15 erg/Gauss, as derived from the measurements of static magnetization curve at 5-30 mg Fe/ml, temperature of about 290K, and in the range of 0-10000 Oersted. PA1 5. The particles have initial magnetic susceptibility of at least about 0.75 EMU/Gauss/g Fe or more as measured in a water-based solution at 5-30 mg Fe/ml and temperature about 290K with applied magnetic field H=50 Gauss. Values of initial magnetic susceptibility up to 1.1-1.2 EMU/Gauss/g Fe are obtainable according to the invention. PA1 6. The particles have specific power absorption rate (SAR) of at least 300, preferably more than 350 W/g Fe as measured in a water-based medium in an electromagnetic field having the frequency of 1 MHz and peak-to-zero amplitude of 100 Oersted. Particles having SAR up to 600 W/g Fe are obtainable according to the invention. PA1 7. The particle hydrodynamic size lies predominantly in the range characterized by sedimentation constant between 150 and 5000 S, preferably between 500 and 2500 S, as determined by centrifugal fractionation in dilute aqueous solution at the temperature of about 290K. PA1 1. time points of 6, 24, 48 and 96 hours are chosen; PA1 2. along with the other tissues as described above, tumor tissue is excised, weighed, and the content of the ferromagnetic iron oxide and the liposome lipid are determined from radioactivity in the similar manner. PA1 (a) Monoclonal antibodies against cancer cell surface antigen are digested by pepsin, and F(ab).sub.2.sup.1 fragments isolated. Specifically, we use KC4G3 murine IgG.sub.3 monoclonal antibody (Coulter) reactive to 400K surface glycoprotein produced by various human adenocarcinomas [74,79] and the F(ab).sup.2 2 is prepared using a commercial reagent kit (Pierce). PA1 (b) Saturated long-chain phophatidylethanolamine is modified by a bifunctional crosslinking reagent to introduce a thiol-binding group. We use distearoylphosphatidylethanolamine (Avanti PolarLipids), and modify it with succinimidyl 4 (-p-maleimidophenyl)butyrate (Pierce) in benzene in the presence of triethylamine as described [70]. The resulting phospholipid derivative, MPB-DSPE carries a thiol-reactive maleimide group at the hydrophilic "head" of its molecule. PA1 (c) F(ab).sup.1 2 fragments are reduced by mercaptoethylamine to give Fab.sup.1 fragments. Fab.sup.1 fragments are purified by gel-filtration using oxygen-free buffer and immediately reacted with MPB-DSPE solubilized by a dialyzable detergent octylglycoside. Since the only available thiol groups are at the hinge region of Fab.sup.1 fragments, the antigen binding sites remain intact. Unreacted maleimide groups are destroyed by mild alkaline treatment (pH 8-8.5). There is no need to separate unreacted Fab.sup.1 fragments at this step since they do integrate into liposomal membrane during further procedures. PA1 (d) PEG-derivatized DSPE, cholesterol, and DSPC or hydrogenated egg phosphatidylcholine (HEPC) are added to the reaction mixture to the proportion established in the previous section. PA1 (e) The reaction mixture is dialyzed against buffer to achieve formation of small unilamellar liposomes (detergent dialysis liposomes) as described [7,374]. These liposomes incorporate both PEG groups and Fab-groups oriented outward from the lipid bilayer. The liposomes are separated from unreacted Fab.sup.1 and residual octylglycoside by gel filtration using Sephacryl S-400 (Pharmacia). PA1 (f) The liposomes obtained at the prior step are used to complete a bilayer formation on the surface of ferrocolloid particles as described in the previous section. The modified ferroliposomes--immunoferroliposomes--are purified from unreacted PEG -and Fab.sup.1 -DSPE liposomes by high gradient magnetic separation as described herein, resuspended in physiological buffer, sterilized by filtration, characterized by their size distribution, lipid and iron content as described in the previous sections, and stored at 4.degree. C. for further studies. PA1 (a) Sonication. Ferrolipid is hydrated with shaking in aqueous buffer (20 mM HEPES, 0.15 M NaCl) or in the solution of a model solute (see below) above Tc. Multilamellar vesicles (MLV) formed in this way are sonicated at the same temperature until clearness (usually 5-6 min). Then the reaction mixture is allowed to cool to room temperature. This method is the simplest and produces small (30-50 nm) unilamellar liposomes. PA1 (b) Solvent injection and sizing through polycarbonate membrane. Ferrolipid is dissolved in a low-boiling nonpolar solvent (methylene dichloride) and injected into water phase as described [47], the temperature being maintained above Tc. The resulting liposome slurry is put under vacuum to remove residual organic solvent and forced through 0.2 .mu.m Nuclepore polycarbonate filter for sizing and separation of liposomes [48]. This method produces so-called large unilamellar/oligolamellar vesicles (LUV) and allows higher loads of a solute. PA1 (c) Rapid extrusion. MLV prepared as above is repeatedly forced through two stacked Nuclepore membranes (0.2 .mu.m) using the syringe extruder similar to the one described in the literature [49]. We have modified the design of the extruder to allow thermostatted operation. Extrusion is be done at a temperature above Tc. Compared to solvent injection, this method allows high loads of the solute without the danger of contamination by an organic solvent [50]; however, it may result in higher losses of iron and/or lipid on the filter. Finally ferroliposomes are passed through Sephadex G-25 desalting column (PD-10, Pharmacia) to remove all low-molecular impurities and stored in refrigerator before use. PA1 a. Size and ultrastructure of ferroliposomes are studied by electron microscopy (EM) as described [49]. Briefly, ferroliposomes are applied on the bacitracin-treated carbon-reinforced grids and contrasted with 1% phosphotungstic acid. Size distribution of ferroliposomes is obtained by direct measurements of EM images. Iron oxide is easily recognized at the electron micrographs as dark electron-dense granules ca. 10 nm in diameter, while contours of the lipid phase are negatively contrasted. PA1 b. Iron and lipid concentration. Iron concentration is determined by o-phenantroline method after digestion of an aliquot with conc. H.sub.2 SO.sub.4 and reduction with ascorbate [51]. Lipid concentration is determined by phosphorous as described [52]. PA1 (a) storage at room temperature and +4.degree. C.; PA1 (b) water bath heating to temperatures above Tc; PA1 (c) exposure to RF electromagnetic field.
4) purification of magnetic particles by gel-filtration.
The heating properties of the resulting ferrocolloids are insufficient to make a qualitative improvement in the amount of ferrocolloid required for heating of a tumor in the radiofrequency electromagnetic (RF EM) fields of such frequency and amplitude that would not at the same time cause substantial background heating of the tissue itself. As explained in numerous publications [3-5, 11,12] such frequencies and amplitudes lie in the range of 0.05-1.2 MHz and 0-200 Oersted, respectively.
The use of magnetic microparticles has been heretofore hindered by the fact that injected particles are rapidly cleared from the bloodstream by reticuloendothelial (RES) cells, primarily in liver and spleen [27, 28, 29]. No successful method of overcoming this problem has been reported to date. The problem of RES-dependent blood clearance is also observed in site-specific drug delivery studies. It has been shown that large unicellular liposomes of 0.1-0.5 .mu.m in size, enriched in sphingomyelin, phospholipids with high transition temperatures and exposed sugar moieties (e.g., phosphatidylinositol or sialogangliosides [30-32] have circulation lifetimes of up to several days, compared to several hours for conventional liposomes [33, 34] and other pharmaceutical microparticles [29]. The mechanism for the extended life of such liposomes is unknown. Possibly reduced opsonization of lipid bilayers can play a role in reduced phagocytosis of the modified liposomes. A method for coating magnetic particles with lipid bilayers has been reported by DeCuyper et al. [35, 36]. Ferrocolloids stabilized by fatty acids are used as a starting material and lipid bilayer from pre-formed small unilamellar liposomes is substituted for the fatty acid micelle. The procedure allows very high loads of ferromagnetic material to be incorporated into the resulting microparticles that have the structure of large unicellular liposomes. Further loading with a drug using liposome fusion is also possible [37].
The therapeutic potential of anticancer drugs is substantially limited by their high systemic toxicity. In attempts to avoid this disadvantage, there have been numerous efforts to deliver anticancer drugs specifically to a tumor. These efforts often used the principle of "recognition" molecules such as antibodies, lectins, desialylated glycoproteins, and autologous tumor lipids [38-40]. However, the modification imposed by a "recognition" molecule, or a bulky construct needed to combine a recognition portion with a drug substance always reduced the activity of the drug and/or its ability for extravasation into tissues, so that only very modest success has been achieved in this field. Another group of studies focused on microspheres as potential targetable drug carriers [41]; however, the yet unsolved problem of rapid reticuloendothelial clearance of such microspheres [27] puts the usefulness of this approach, in its present state, under question.
A different approach to drug targeting has been developed in the works by Yatvin et al. [42,43] and Huang et al. [44]. They used heat to induce rapid release of pharmaceuticals from thermosensitive liposomes composed of phospholipids having transition temperatures slightly above normal physiological temperature. Local hyperthermia, heating of the target area to a temperature of 42.degree.-44.degree. C., would cause the liposome lipids to "melt", and the liposomes flowing through the vascular bed of a hyperthermized area would rapidly release the entrapped drug into the surrounding medium. Since the drug is released in its intact form, the problems concerning drug extravasation and activity are avoided. So, in the approaches proposed by Yatvin and Huang, the targeted mode of drug delivery substantially depends on the ability to apply hyperthermia to the area of pathology in a targeted manner; unfortunately, none of the existing techniques of hyperthermia offers a general and satisfactory way to do so [10].
In the technique of magnetic resonance imaging (MRI), colloidal gamma ferric oxide has been used in human patients as a contrast enhancer [17]. The compound is virtually non-toxic and well tolerated when administered into the bloodstream. Its LD.sub.50 in mice exceeds 2 g per kg [5]. No toxic effects were observed either upon acute (150 mg Fe/kg in 24 hours) or chronic administration in rats [14] or in human patient trials [15]. A principal magnetic property of importance for particles used in MRI imaging is the effect on the T.sub.2 proton relaxation time. Contrast enhancement is increased by agents which decrease the T.sub.2 relaxation time. (See, e.g., Cerdan, S., et al., (1989) Magnetic Resonance in Medicine, 12:151-163.)